Biodegradable polymers have been used to deliver various therapeutic agents. The therapeutic agents typically are encapsulated within the biodegradable polymers which are formed into particles having sizes of 100 μm or less, films, sheets, disks, pellets, or implants. The biodegradable polymers are administered to a person, and the encapsulated therapeutic agent is released within the body of the patient as the polymer degrades and/or as water diffuses into the polymer to leach out the encapsulated therapeutic. Biodegradable polymers, both synthetic and natural, can release encapsulated agents over a period of days or weeks, which can have benefits in administration of drugs or other agents.
These devices have been modified to incorporate drug through such techniques as solvent encapsulation, melt encapsulation, phase separation, and other standard methods for processing of polymers. The surfaces of the polymeric devices have been modified to incorporate ligands, usually through either derivatization of the polymer before formation of the device, or after formation of the device using covalent binding to the polymer or ionic binding to charged sites on the polymer. Many of these techniques have disadvantages. Derivatization of the polymer prior to formation of the device can result in many of the ligands being encapsulated within the device, lowering the useful number of ligands available for binding or targeting. Covalent binding after formation can damage the polymers, lead to cross-reactions that decrease specificity, and is typically not highly efficient. Ionic binding is very gentle, but subject to dissociation, frequently not possible in high density, and of low specificity.
Biodegradable polymers fabricated from poly(lactic-co-glycolic acid) (PLGA) have emerged as powerful potential carriers for small and large molecules of therapeutic importance as well as scaffolds for tissue engineering applications. This importance derives from: 1) Physiologic compatibility of PLGA and its hompolymers PGA and PLA, all of which have been established as safe in humans after 30 years in various biomedical applications including drug delivery systems 2) Commercial availability of a variety of PLGA formulations for control over the rate and duration of molecules released for optimal physiological response (Visscher et al. J Biomed Mater Res 1985; 19(3):349-65; Langer R, Folkman J. Nature 1976; 263(5580):797-800; Yamaguchi. J. Controlled Rel. 1993; 24(1-3):81-93.). 3) Biodegradability of PLGA materials, which provides for sustained release of the encapsulated molecules under physiologic conditions while degrading to nontoxic, low-molecular-weight products that are readily eliminated (Shive et al. Adv Drug Deliv Rev 1997; 28(1):5-24; Johansen et al. Eur J Pharm Biopharm 2000; 50(1):129-46). 4) Control over its manufacturing into nanoscale particles (<500 nm) for potential evasion of the immune phagocytic system or fabrication into microparticles on the length scale of cells for targeted delivery of drugs or as antigen-presenting systems (Eniola et al. J Control Release 2003; 87(1-3):15-22; Jain R A. Biomaterials 2000; 21(23):2475-90). This unique combination of properties coupled with flexibility over fabrication has led to interest in modifying the PLGA surface for specific attachment to cells or organs in the body (Eniola, et al. 2003; Keegan et al., Biomaterials 2003; 24(24):4435-4443; Lamprecht et al. J Pharmacol Exp Ther 2001; 299(2):775-81; Lathia et al. Ultrasonics 2004; 42(1-9):763-8 Park et al. J Biomed Mater Res 2003; 67A(3):751-60; Panyam Adv Drug Deliv Rev 2003; 55(3):329-47) for drug delivery and tissue engineering applications. With a functional PLGA surface, cells may be attached specifically to scaffolds enabling control over interactions that lead to formation of optimal neotissue, or encapsulated drug or antigen delivered specifically to the site of interest potentially reducing deleterious drug side effects and enhancing antigen delivery for vaccine applications.
A major difficulty associated with coupling ligands to PLGA particles has been the lack of functional chemical groups on the aliphatic polyester backbone for linking to target ligands. This severely hinders the application of traditional conjugation methods to the PLGA surface. Thus to introduce functionality into PLGA surfaces several approaches have been studied. These include, synthesis of PLGA copolymers with amine (Lavik et al J Biomed Mater Res 2001; 58(3):291-4; Caponetti et al. J Pharm Sci 1999; 88(1):136-41) or acid (Caponetti et al J Pharm Sci 1999; 88(1):136-41) end groups followed by fabrication into particles. Another approach involves the blending or adsorption of functional polymers such as polylysine (Faraasen et al. Pharm Res 2003; 20(2):237-46; Zheng et al. Biotechnology Progress 1999; 15(4):763-767) or poly(ethylene-alt-maleic acid) (PEMA)(Keegan et al. Macromolecules 2004) or PEG (Muller J Biomed Mater Res 2003; 66A(1):55-61) into PLGA and forming particles and matrices from these blends (Zheng, et al. 1999; Keegan, 2004; Park et al. J Biomater Sci Polym Ed 1998; 9(2):89-110; Croll Biomacromolecules 2004; 5(2):463-73; Cao et al. Methods Mol Biol 2004; 238:87-112). Plasma treatment of the PLGA matrix has also been proposed for the purpose of modifying its surface properties and introducing hydrophilic functional groups into the polymer (Yang et al. J Biomed Mater Res 2003; 67A(4):1139-47; Wan et al., Biomaterials 2004; 25(19):4777-83).
Targeting ligands include any molecule that recognizes and binds to target antigen or receptors over-expressed or selectively expressed by particular cells or tissue components. These may include antibodies or their fragments, peptides, glycoproteins, carbohydrates or synthetic polymers. The most widely used coupling group is poly(ethylene glycol) (PEG), because this group creates a hydrophilic surface that facilitates long circulation of the nanoparticles. This strategy has been used successfully in making ‘Stealth’ liposomes with affinity towards target cells. Incorporating ligands in liposomes is easily achieved by conjugation to the phospholipid head group, in most cases phosphotidylethanolamine (PE), and the strategy relies either on a preinsertion of the functionalized lipid or post insertion into a formed liposome. Functionality could also be introduced by incorporating PEG with functional endgroups for coupling to target ligands.
While these approaches have had good success in their specific applications, their general use is hindered by drawbacks such as difficulty associated with preparing the needed copolymers, limited density of functional groups and targeting effects that decrease with time due to desorption or degradation of adsorbed group as the particle or scaffold erodes. It would be most desirable to retain ligand function with control over its density on the surface for prolonged periods of time for improved drug delivery. There are also still a number of difficulties associated with preparation of co-polymers, limited density of functional groups and targeting groups with time due to degradation.
It is therefore an object of the present invention to provide a polymer delivery system which can preferentially deliver therapeutic compositions to selected cells or tissue and/or deliver high amounts of therapeutic molecules.
It is another object of the invention to provide high density, direct attachment to polymer, without harsh cross-linking or coating requirements.